X-ray Apparatus

ABSTRACT

X-ray apparatus comprises a linear accelerator adapted to produce a beam of electrons at one of at least two selectable energies and being controlled to change the selected energy on a periodic basis, and a target to which the beam is directed thereby to produce a beam of x-radiation, the target being non-homogenous and being driven to move periodically in synchrony with the change of the selected energy. In this way, the target can move so that a different part is exposed to the electron beam when different pulses arrive. This enables the appropriate target material to be employed depending on the selected energy. The easiest form of periodic movement for the target is likely to be a rotational movement. The target can be immersed in a coolant fluid such as water. The linear accelerator can be of the type disclosed in WO2006/097697A1. The target preferably contains at least one exposed area of tungsten and/or at least one exposed area of carbon. These can be present as inhomogeneities in the material of which the target is composed, such as Carbon inserts in a Tungsten substrate (or vice versa), alternating segments of Carbon and Tungsten, Carbon and Tungsten inserts in a substrate of a third material, or arrangements involving other materials in addition to or instead of Carbon and/or Tungsten. Alternatively, the target can be of a homogenous material but have inhomogeneities in its thickness to cater for the different electron energies. The same concept can be applied to the filter. A detector can be provided, operating in synchrony with the energy variation. Such an x-ray apparatus can form a part of a radiotherapy apparatus, in which case the first selected energy can be a diagnostic energy and a second selected energy a therapeutic energy.

FIELD OF THE INVENTION

The present invention relates to x-ray apparatus.

BACKGROUND ART

In the use of radiotherapy to treat cancer and other ailments, apowerful beam of the appropriate radiation is directed at the area ofthe patient that is affected. This beam is apt to kill living cells inits path, hence its use against cancerous cells, and therefore it ishighly desirable to ensure that the beam is correctly aimed. Failure todo so may result in the unnecessary destruction of healthy cells of thepatient.

Several methods are used to check this, and devices such as the Elekta™Synergy™ device employ two sources of radiation, a high energyaccelerator capable of creating a therapeutic beam and a lower energyX-ray tube for producing a diagnostic beam. Both are mounted on the samerotateable gantry, separated by 90°. Each has an associated flat-paneldetector, for portal images and diagnostic images respectively.

In our earlier application WO-A-99/40759, we described a novel couplingcell for a linear accelerator that allowed the energy of the beamproduced to be varied more easily than had hitherto been possible. Inour subsequent application WO-A-01/11928 we described how that structurecould be used to produce very low energy beams, suitable for diagnosticuse, in an accelerator that was also able to produce high-energytherapeutic beams. Later, in WO2006/097697A1 we described how to switchbetween those high- and low-energy beams at high speed. The disclosureof all of these prior disclosures is hereby incorporated by reference.The reader should note that this application develops the principles setout in those applications, which should therefore be read in conjunctionwith this application and whose disclosure should be taken to form partof the disclosure of this application.

SUMMARY OF THE INVENTION

The Elekta™ Synergy™ arrangement works very well, but requires someduplication of parts in that, in effect, the structure is repeated toobtain the diagnostic image. In addition, care must be taken to ensurethat the two sources are in alignment so that the diagnostic view can becorrelated with the therapeutic beam. However, this has been seen asnecessary so that diagnostic images can be acquired during treatment toensure that the treatment is proceeding to plan.

WO-A-01/11928 shows how the accelerator can be adjusted to produce alow-energy beam instead of a high-energy beam, and WO2006/097697 A1shows how the two beams could be produced (effectively) simultaneouslyas is required for concurrent therapy and monitoring.

The present invention therefore provides an X-ray apparatus comprising alinear accelerator adapted to produce a beam of electrons at one of atleast two selectable energies and being controlled to change theselected energy on a periodic basis, and a target to which the beam isdirected thereby to produce a beam of x-radiation, the target beingnon-homogenous and being driven to move periodically in synchrony withthe change of the selected energy.

In this way, the target can move so that a different part is exposed tothe electron beam when different pulses arrive. This enables theappropriate target material to be employed depending on the selectedenergy.

The easiest form of periodic movement for the target is likely to be arotational movement. The target can be immersed in a coolant fluid suchas water.

The linear accelerator can be of the type comprising a series ofaccelerating cavities, adjacent pairs of which are coupled via couplingcavities, at least one coupling cavity comprising a rotationallyasymmetric element that is rotateable thereby to vary the couplingoffered by that cavity and thereby select an energy. It can furthercomprise a control means adapted to control operation thereof andcontrol rotation of the asymmetric element, arranged to operate theaccelerator in a pulsed manner and to rotate the asymmetric elementbetween pulses to control the energy of successive pulses. Generally, weprefer that rotation of the asymmetric element is continuous duringoperation of the linear accelerator.

The target preferably contains at least one exposed area of a firstmaterial and/or at least one exposed area of a second material. Suitablematerials are tungsten and carbon, but others will also be suitable.These can be present as inhomogeneities in the material of which thetarget is composed, such as Carbon inserts in a Tungsten substrate (orvice versa), alternating segments of Carbon and Tungsten, Carbon andTungsten inserts in a substrate of a third material, or arrangementsinvolving other materials in addition to or instead of Carbon and/orTungsten.

Alternatively, or in addition, the target can have inhomogeneities inits thickness to cater for the different electron energies. Thicknessdifferences may cause interesting weight distributions (depending ontheir spatial distribution), which could be balanced by partially, fullyor over-filling the thinner areas with an inert material.

Most X-ray apparatus include one or more filters for the x-radiation,such as flattening filters and diagnostic x-ray filters. These areusually matched to the energy distribution of the x-rays being filtered.We therefore propose that the apparatus comprise a filter housing, inwhich there are a plurality of filters, the housing being driven to moveperiodically in synchrony with the change of the selected energy, i.e. afilter using essentially the same inventive concept as that set outabove in relation to the target.

Accordingly, the present invention further provides an X-ray apparatuscomprising a linear accelerator adapted to produce a beam of electronsat one of at least two selectable energies and being controlled tochange the selected energy on a periodic basis, a target to which thebeam is directed thereby to produce a beam of x-radiation, and a filterhousing, in which there are a plurality of filters for the x-radiation,the housing being driven to move periodically in synchrony with thechange of the selected energy.

A detector can be located in the path of the beam, to acquire an imageproduced by the beam after attenuation thereof. This is preferablydriven by a control means operating in synchrony with the control ofchanges to the selected energy of the linear accelerator.

The above x-ray apparatus can, for example, form a part of aradiotherapy apparatus. In that case, the first selected energy can be adiagnostic energy and a second selected energy a therapeutic energy.

BRIEF DESCRIPTION OF THE DRAWINGS

An embodiment of the present invention will now be described by way ofexample, with reference to the accompanying figures in which;

FIG. 1 shows a view of a pair of accelerator cavities and the couplingcavity between them;

FIGS. 2 and 3 show characteristic curves for the accelerator, FIG. 2showing the variation in linac impedance with vane angle;

FIG. 4 shows an arrangement for rotating the asymmetric element;

FIG. 5 shows an axial section along an x-ray apparatus according to thepresent invention; and

FIGS. 6 to 11 show alternative designs of target for the x-ray apparatusof FIG. 5.

DETAILED DESCRIPTION OF THE EMBODIMENTS

Our application WO2006/097697A1 showed the basis of an x-ray apparatusable to switch effectively ‘instantaneously’ from a therapeutic energyto an imaging energy, to allow imaging during therapy but with nooverhead in time and utilising a much simpler construction. FIG. 1 showsthe coupling cavity of the linac 10 disclosed in WO-A-99/40759 andWO2006/097697A1. A beam 12 passes from an ‘n^(th)’ accelerating cavity14 to an ‘n+1^(th)’ cavity 16 via an axial aperture 18 between the twocavities. Each cavity also has a half-aperture 18 a and 18 b so thatwhen a plurality of such structures are stacked together, a linearaccelerator is produced.

Each adjacent pair of accelerating cavities can also communicate via“coupling cavities” that allow the radiofrequency signal to betransmitted along the linac and thus create the standing wave thataccelerates electrons. The shape and configuration of the couplingcavities affects the strength and phase of the coupling. The couplingcavity 20 between the n^(th) and n+l^(th) cavities is adjustable, in themanner described in WO-A-99/40759, in that it comprises a cylindricalcavity in which is disposed a rotateable vane 22. As described inWO-A-99/40759 and WO-A-01/11928 (to which the skilled reader isreferred), this allows the strength and phase of the coupling betweenthe accelerating cells to be varied by rotating the vane, as a result ofthe rotational asymmetry thereof.

It should be noted that the vane is rotationally asymmetric in that asmall rotation thereof will result in a new and non-congruent shape tothe coupling cavity as “seen” by the rf signal. A half-rotation of 180°will result in a congruent shape, and thus the vane has a certain degreeof rotational symmetry. However, lesser rotations will affect couplingand therefore the vane does not have complete rotational symmetry; forthe purposes of this invention it is therefore asymmetric.

The n^(th) accelerating cavity 14 is coupled to the n−1^(th) by a fixedcoupling cell. That is present in the structure illustrated in FIG. 1 asa half-cell 24. This mates with a corresponding half-cell in theadjacent structure. Likewise, the n+1^(th) accelerating cell 16 iscoupled to the n+2^(th) such cell by a cell made up of the half-cell 26and a corresponding half-cell in an adjacent structure.

The radiation is typically produced from the linac in short pulses ofabout 3 microseconds, approximately every 2.5 ms. To change the energyof a known linac, be that by way of the rotateable vane described aboveor by other previously known means, the linac is switched off, thenecessary adjustment is made, and the linac is re-started.

According to the invention, the rotateable vane 22 is caused tocontinuously rotate with a period correlated to the pulse rate of thelinac. Thus, in this example the period is 2.5 ms i.e. 400 revolutionsper second or 24,000 rpm. The radiation is then produced at a particularposition of the vane or a particular phase of the rotation. Given thatthe linac is active for only 0.12% of the time, the vane will (at most)rotate through slightly less than half a degree and thus will bevirtually stationary as “seen” by the rf signal.

This phase of the linac's pulse can be easily changed from one pulse tothe next. This therefore allows the energy to be switched from one pulseto the next, since changing the phase correlates with the selection of adifferent vane angle.

In the adjustable coupling cell 20, the electric fields are symmetricalon either side of the vane. It therefore follows that the vane spinspeed can in fact be reduced by a factor of 2 compared to that suggestedabove, which allows a lesser spin speed of 12,000 rpm to be adopted.

FIG. 2 illustrates a practical aspect of the use of such a system. Asmay be seen in the Voltage Standing Wave Ratio (VSWR) vs vane angleplot, there are two “danger zones” in the angle ranges of 100°-120° and280°-300°, in which the waveguide is under coupled. They should beavoided, by use of a suitable control mechanism.

Within the working range of 120° to 280°, there are benefits inadjusting the input power according to the vane angle, to maintain theelectric field constant. This is mainly due to the fact that the VSWR ofthe whole waveguide changes with the vane angle. FIG. 3 shows the inputpower required (in brackets) at different angles, together with thevarying electrical field developed after the adjustable coupling cell at200 mm along the linac. These varying electric fields translate into avarying energy of the electrons produced by the linac. Note that at 264°the electric field after the adjustable coupling cell is reversed; thisdecelerates the electrons and results in a very low diagnostic energy asdescribed in WO-A-01/11928.

This idea can also be used to servo the actual energy of the beam totake account of variations in other systems.

The ability to vary the energy pulse to pulse could be used to controlthe depth dose profile pulse to pulse. This could be of benefit on ascanned beam machine where the ability to vary the energy across theradiation field could be used to produce less rounded isodose lines.

A further advantage of being able to vary the energy so rapidly would beto vary the therapy beam energy when in electron mode, thereby extendingthe irradiated volume receiving 100% of the dose. This could also beuseful in Energy modulated electron therapy (EMET) or modulated electronradiotherapy (MERT) techniques. The fast switching of the electronenergy and possibly the scattering foil could enable these techniques tobe delivered more quickly, provided that suitable electron beamcollimation could be provided.

FIG. 4 shows a possible mechanism by which the vane 22 can be rotatedcontinuously. The vane does of course sit in an evacuated volume, soevidently a suitable shaft could be provided, with appropriate sealing,to transmit rotation from a motor outside the evacuated volume.Alternatively, as shown illustratively in FIG. 4, a magnetic controlsystem could be provided. In this arrangement, the vane 22 is providedwith magnetically polarised sections 28, 30 on either end. Then, outsidethe vacuum seal 32, an array of electrical coils 34, 36 etc areprovided. These can then interact with the polarised sections 28, 30 inthe manner of a stepper motor.

The above description allows for the production of a beam of electronsat a selectable energy. This can then be converted to a beam of x-raysby directing the electron beam at a suitable target. According to knownprinciples of x-ray production, this produces a beam of x-rays which canthen be collimated (etc) to produce a therapeutically or diagnosticallyuseful output.

A potential problem in this is that the target is usually chosen in thelight of the electron and x-ray energies involved. For example, a lowerenergy diagnostic beam (i.e. one comprising low energy photons such aswith an energy below 200 KeV) can be produced from a megavoltageelectron beam by directing the beam to a thinner or a lower atomicnumber target, Carbon being one example (see D. M. Galbraith,“Low-energy imaging with high-energy bremsstrahlung beams”, Med. Phys.16(5), 734-46 (1989)), whereas a high energy therapeutic beam isproduced by directing a suitable electron beam to a thicker or higheratomic number target, Tungsten being an example. Whilst it is possibleto select a compromise target material, a better beam quality isachievable by matching the target material to the selected energy.

In fact, in such circumstances, the Carbon target serves two purposes—toproduce photons and to remove electrons which would otherwise increasethe patient skin dose. At very low energies (circa 400 KeV) the majorityof photons can arise from the electron window itself, and thus asignificant part of the function of the Carbon target is to act as anelectron filter.

This can be done as shown in FIG. 5. A linear accelerator comprises aseries of sequential accelerating cells 102, 104, 106, 108 etc. Betweencells 106 and 108, the third and fourth cells, there is a variablecoupling cell 110 which is designed according to the principles of thevariable coupling cell 20 of FIG. 1 and includes a continuously rotatingvane 112 as described with respect to FIG. 4. The accelerator isenclosed within a vacuum enclosure 114 which has an output window 116through which the electron beam produced by the linear accelerator 100passes. The beam then impinges on a target 118.

The target 118 is generally disc-shaped and is mounted on a central axle120 which is driven by an external motor (not shown) so that the target118 rotates. The target 118 and the axle 120 are located relative to thelinear accelerator 100 so that the electron beam impinges at a locationon the target that is offset from the centrally-mounted axle 120. Thus,as the target 118 rotates, the relatively narrow electron beam will passthrough the disc-shaped target at a point or points on a circular path.

The target 118 is rotationally asymmetric, and includes differentregions made up of different materials. Thus, as the electron beamimpinges on different parts of the target 118, a different targetmaterial is presented at the point of impingement. It therefore onlyremains to control the rotation and/or the pulse timings so thatsuccessive pulses of differing energy electron beams meet theappropriate location on the target 118.

FIGS. 6 to 11 show different possible designs of the target 118. FIG. 6shows a simple target 122 that is constructed from two half-discs 124,126, each semicircular in plan view. In this example, one is of Tungstenand the other is of Carbon, and the two are joined along their straightedge to form a single disc-shaped target 122. As this rotates, italternately presents W or C locations to the impinging electron beam128. Provided that rotation of the target 122 is synchronised to thevarying energy pulses, the appropriate target material will therefore bepresented at the appropriate time.

FIG. 7 shows an alternative design of target 130. Instead of beingdivided into halves, this target 130 is divided into four quarters.Alternate quarters are of alternating material, thus as the target 130rotates, the path 132 followed by the electron beam across the target130 traverses a Tungsten quarter 134, which is then replaced by a Carbonquarter 136, then by a Tungsten quarter 138, then by a Carbon quarter140 which is then replaced by the original Tungsten quarter 134 after acomplete revolution. At the expense of a slight increase inconstructional complexity, the permits the rotational speed of thetarget to be halved.

Naturally, a greater number of segments could be provided in order topermit the rotational speed to be reduced still further. Even numberssuch as 6, 8, 10 segments (etc) will suit arrangements in which twotarget materials are provided, but other numbers may be suited toarrangements using three or more different target materials, or thetarget geometry could be adjusted in this way to cater for periodicvariations in pulse timing. For example, if the variation in outputenergy is used to control the depth penetration of the radiation thenprovision might be made for an option to provide an occasional pulse ata different position of the rotating vane 112 in order to allow such athird energy level. This would be at a different phase point, and couldthus be made to correspond to a different segment of the target.

FIG. 8 shows a further form of target 142 in which a larger Tungstenarea 144 and a smaller Carbon area 146 are joined to form thedisc-shaped target 142. Thus, the join between the two segments is anacute angle, with the larger Tungsten segment occupying about 240° andthe smaller Carbon segment being the remainder. The path 150 traced onthe target 142 by the electron beam thus spends longer on the Tungstensegment 144; this could be useful if the therapeutic beam energy is tobe varied, as this will necessitate waiting for a slightly differentposition of the rotating vane 112 and hence a different phase point; thegreater area of the Tungsten segment 144 allows some latitude toaccommodate this variation in timing. Of course, a larger Carbon segmentcould alternatively be provided if multiple diagnostic energies are tobe provided, as is sometimes called for.

FIG. 9 shows a potentially more robust target 152 in which a smallerdisc 154 of Carbon is inset within a suitable aperture in a larger disc156 of Tungsten. As the target 152 rotates, the Carbon disc 154 isretained more securely in the Tungsten disc 156, whilst the path 158traced by the electron beam still alternates between Carbon andTungsten. The materials could of course be reversed as required.

FIG. 10 shows a slower-rotating version 160 of the target of FIG. 9. ATungsten disc 162 has several apertures, in this case three, in whichCarbon discs 164, 166, 168 are placed. Thus, as the target 160 rotates,the path 170 of the electron beam again alternates between Tungsten andCarbon but does so several times in one revolution. Accordingly, therotational velocity can be reduced. Naturally, a greater or lessernumber of inserts 164, 166, 168 can be provided as desired, and/or thematerials reversed.

FIG. 11 shows a slightly different design of target 172. A substrate 174is generally disc-shaped, and can be of any material having suitablemechanical properties. Two generally semi-circular inserts 176, 178 areprovided in the substrate 174, one of Tungsten and the other of Carbon.As the target 172 rotates, the path 180 traced by the electron beamcrosses alternately from the Tungsten insert 176 to the Carbon insert178. As the beam path crosses from one to the other, it briefly passesover the substrate material, but it is to be expected that the pulsetiming will be adjusted so that such “crossover” times are not chosenfor a pulse, as minor errors in the pulse timing may result inmisplacing the beam.

Other geometries for the inserts could be adopted, following the generalgeometries of FIGS. 6 to 11, or otherwise. Likewise, other rotationallyasymmetric geometries for the targets of FIGS. 6 to 11 could be adopted.

It should be emphasised that other materials could be used for theactive regions of the targets. Tungsten and Carbon have been used in theabove discussion as examples as they are the most common choices, butother materials are also suitable.

Returning to FIG. 5, the x-ray beam 182 produced at the rotating target118 is then limited generally by a primary collimator 184. Normally, thebeam will be filtered at this point, such as to flatten it or fordiagnostic purposes. Diagnostic x-ray filters are usually made ofAluminium and enable the quality of the x-ray beam to be adjusted, forexample to remove very low energy photons (<30 KeV) from an x-ray beamand thereby reduce the patient skin dose. Again, the filter willtypically be specific to the beam energy, presenting a potentialdifficulty if the beam energy varies.

Thus, a flattening filter can be omitted or replaced with a uniformmaterial and an unflattened beam employed (according to generally knownprinciples).

Alternatively (as illustrated) a rotating filter housing 186 can beprovided. This is a disc-shaped substrate carrying a plurality offilters, usually two, located in the substrate at an angular position sothat when a pulse of a specific energy is emitted from the target 118,the appropriate filter is presented by the rotating filter substrate186. If a flattening filter is used in this housing, then it is requiredthat it is accurately positioned. Using an unflattened beam has theadvantage of using a uniform or no filter for which the position is notcritical.

From there, the beam then passes through an ion chamber 188, amulti-leaf collimator 190 and a block collimator 192, and/or suchcollimation as is required for the specific application in which thex-ray apparatus is employed. FIG. 5 also shows a mirror 194 placed inthe path of the beam 182; this can be used to project visible light froma lamp 196 and filter 198 along the beam path 182 and hence checkalignment, patient positioning etc.

Some form of detector will be needed for at least the diagnosticradiation. A range of flat panel detectors are suitable, and many areable to withstand the higher energy therapeutic radiation that will betransmitted through the patient. In particular, GEM (Gas ElectronMultiplier) detectors, solid state, and CCD detectors, and active pixelsensors based on CMOS technology could be suitable and at least one canbe located on the beam path with the patient between it and theapparatus shown in FIG. 5.

A suitable detector could be based on the technologies illustrated anddescribed in U.S. Pat. No. 6,429,578 B1, WO 2005/120046, and EP1762088,in the thesis “New Efficient Detector for Radiation Therapy Imagingusing Gas Electron Multipliers” submitted by Janina Ostling to StockholmUniversity, 17 Mar. 2006, ISBN 91-7155-218-9, and in the paper“Empirical electro-optical and X-ray performance evaluation of CMOSactive pixels sensor for low dose, high resolution X-ray medicalimaging” by Costas Arvanitis, Sarah Bohndiek, Gary Royle, Andrew Blue,Huang XingLiang, Andy Clark, Mark Prydderch, Renato Turchetta, andRobert Speller, Medical Physics 34 (2007) 4612-4625. Active pixelsensors are discussed in the article available athttp://medicalphysicsweb.org/cws/article/research/31467. The contents ofthese documents are incorporated herein by reference, and the readershould be aware that the present application should be read inconjunction with these documents, the content of which may be used byway of amendment to this application.

The detector of this example is operated in synchrony with the switchingenergy. To capture images from the low energy pulse only, the detectorcan be reset immediately after a high energy pulse. Alternatively, tocapture both low energy images and portal images, the detector can beswitched between modes adapted to each energy in synchrony with theenergy switching.

It will of course be understood that many variations may be made to theabove-described embodiment without departing from the scope of thepresent invention.

1. X-ray apparatus comprising a linear accelerator adapted to produce abeam of electrons at one of at least two selectable energies and beingcontrolled to change the selected energy on a periodic basis, and atarget to which the beam is directed thereby to produce a beam ofx-radiation, the target being non-homogenous and being driven to moveperiodically in synchrony with the change of the selected energy. 2.X-ray apparatus according to claim 1 in which the target movesrotationally.
 3. X-ray apparatus according to claim 1 in which thelinear accelerator comprises a series of accelerating cavities, adjacentpairs of which are coupled via coupling cavities, at least one couplingcavity comprising a rotationally asymmetric element that is rotateablethereby to vary the coupling offered by that cavity and thereby selectan energy.
 4. X-ray apparatus according to claim 3, further comprising acontrol means for the linear accelerator adapted to control operationthereof and control rotation of the asymmetric element, the controlmeans being arranged to operate the accelerator in a pulsed manner andto rotate the asymmetric element between pulses to control the energy ofsuccessive pulses.
 5. X-ray apparatus according to claim 3 in whichrotation of the asymmetric element is continuous during operation of thelinear accelerator.
 6. X-ray apparatus according to claim 1 in which thetarget is immersed in a coolant fluid.
 7. X-ray apparatus according toclaim 6 in which the fluid is predominantly water.
 8. X-ray apparatusaccording to claim 1 in which the target contains at least one exposedarea of a first material and at least one exposed area of a secondmaterial that is different to the first material.
 9. X-ray apparatusaccording to claim 1 in which the target contains at least one exposedarea of tungsten.
 10. X-ray apparatus according to claim 1 in which thetarget contains at least one exposed area of carbon.
 11. X-ray apparatusaccording to claim 1 in which the target has inhomogeneities in itsthickness.
 12. X-ray apparatus according to claim 1 in which the targetis of a homogenous material but has inhomogeneities in its thickness.13. X-ray apparatus according to claim 1 including a filter housing, inwhich there is a plurality of filters for the x-radiation, the housingbeing driven to move periodically in synchrony with the change of theselected energy.
 14. X-ray apparatus comprising a linear acceleratoradapted to produce a beam of electrons at one of at least two selectableenergies and being controlled to change the selected energy on aperiodic basis, a target to which the beam is directed thereby toproduce a beam of x-radiation, and a filter housing, in which there area plurality of filters for the x-radiation, the housing being driven tomove periodically in synchrony with the change of the selected energy.15. X-ray apparatus according to claim 14 comprising a detector locatedin the path of the beam to acquire an image produced by the beam afterattenuation thereof.
 16. X-ray apparatus according to claim 15 in whichthe detector is driven by a control means therefor operating insynchrony with the control of changes to the selected energy of thelinear accelerator.
 17. Radiotherapy apparatus comprising a source ofX-radiation according to claim
 14. 18. Radiotherapy apparatus accordingto claim 17 in which a first selected energy is a diagnostic energy anda second selected energy is a therapeutic energy. 19-20. (canceled) 21.X-ray apparatus according to claim 1 comprising a detector located inthe path of the beam to acquire an image produced by the beam afterattenuation thereof.
 22. Radiotherapy apparatus comprising a source ofX-radiation according to claim 1.